Breathing gas delivery method and apparatus

ABSTRACT

An improved methodology and systems for delivery of breathing gas such as for the treatment of obstructive sleep apnea through application of alternating high and low level positive airway pressure within the airway of the patient with the high and low airway pressure being coordinated with the spontaneous respiration of the patient, and improved methods and apparatus for triggering and for leak management in such systems.

BACKGROUND OF THE INVENTION

[0001] The sleep apnea syndrome, and in particular obstructive sleepapnea, afflicts an estimated 4% to 9% of the general population and idue to episodic upper airway obstruction during sleep. Those afflictedwith obstructive sleep apnea experience sleep fragmentation andintermittent, complete or nearly complete cessation of ventilationduring sleep with potentially severe degrees of oxyhemoglobinunsaturation. These features may be translated clinically intodebilitating daytime sleepiness, cardiac disrhythmias, pulmonary-arteryhypertension, congestive heart failure and cognitive dysfunction. Othersequelae of sleep apnea include right ventricular dysfunction with corpulmonale, carbon dioxide retention during wakefulness as well as duringsleep, and continuous reduced arterial oxygen tension. Hypersomnolentsleep apnea patients may be at risk for excessive mortality from thesefactors as well as from an elevated risk for accidents such as whiledriving or operating other potentially dangerous equipment.

[0002] Although details of the pathogenesis of upper airway obstructionin sleep apnea patients have not been fully defined, it is generallyaccepted that the mechanism includes either anatomic or functionalabnormalities of the upper airway which result in increased air flowresistance. Such abnormalities may include narrowing of the upper airwaydue to suction forces evolved during inspiration, the effect of gravitypulling the tongue back to appose the pharyngeal wall, and/orinsufficient muscle tone in the upper airway dilator muscles. It hasalso been hypothesized that a mechanism responsible for the knownassociation between obesity and sleep apnea is excessive soft tissue inthe anterior and lateral neck which applies sufficient pressure oninternal structures to narrow the airway.

[0003] The treatment of sleep apnea has included such surgicalinterventions as uvalopalatopharyngoplasty, gastric surgery for obesity,and maxillo-facial reconstruction. Another mode of surgical interventionused in the treatment of sleep apnea is tracheostomy. These treatmentsconstitute major undertakings with considerable risk of post-operativemorbidity if not mortality. Pharmacologic therapy has in general beendisappointing, especially in patients with more than mild sleep apnea.In addition, side effects from the pharmacologic agents that-have beenused are frequent. Thus, medical practitioners continue to seeknon-invasive modes of treatment for sleep apnea with high success ratesand high patient compliance including, for example in cases relating toobesity, weight loss through a regimen of exercise and regulated diet.

[0004] Recent work in the treatment of sleep apnea has included the useof continuous positive airway pressure (CPAP) to maintain the airway ofthe patient in a continuously open state during sleep. For example, U.S.Pat. No. 4,655,213 and Australian patent AU-B-83901/82 both disclosesleep apnea treatments based on continuous positive airway pressureapplied within the airway of the patient.

[0005] Also of interest is U.S. Pat. No. 4,773,411 which discloses amethod and apparatus for ventilatory treatment characterized as airwaypressure release ventilation and which provides a substantially constantelevated airway pressure with periodic short term reductions of theelevated airway pressure to a pressure magnitude no less than ambientatmospheric pressure.

[0006] Publications pertaining to the application of CPAP in treatmentof sleep apnea include the following:

[0007] 1. Lindsay, D A, Issa F G, and Sullivan C. E. “Mechanisms ofSleep Desaturation in Chronic Airflow Limitation Studied with NasalContinuous Positive Airway Pressure (CPAP)”, Am Rev Respir Dis, 1982;125: p.112.

[0008] 2. Sanders M H, Moore S E, Eveslage J. “CPAP via nasal mask: Atreatment for occlusive sleep apnea”, Chest, 1983; 83: pp. 144-145.

[0009] 3. Sullivan C E, Berthon-Jones M, Issa F G. “Remission of severeobesity-hypoventilation syndrome after short-term treatment during sleepwith continuous positive airway pressure”, Am Rev Respir Dis, 1983; 128:pp. 177-181.

[0010] 4. Sullivan C E, Issa F G, Berthon-Jones M, Eveslage. “Reversalof obstructive sleep apnea by continuous positive airway pressureapplied through the nares”, Lancet, 1981; 1: pp. 862-865.

[0011] 5. Sullivan C E, Berthon-Jones M. Issa F G. “Treatment ofobstructive apnea with continuous positive airway pressure appliedthrough the nose”, Am Rev Respir Dis, 1982; 125: p.107. Annual MeetingAbstracts.

[0012] 6. Rapoport D M, Sorkin B, Garay S M, Goldring R M. “Reversal ofthe ‘Pickwickian Syndrome’ by long-term use of nocturnal nasal-airwaypressure”, N Engl J. Med, 1982; 307: pp.931-933.

[0013] 7. Sanders M H, Holzer B C, Pennock B E. “The effect of nasalCPAP on various sleep apnea patterns”, Chest, 1983; 84: p.336. Presentedat the Annual Meeting of the American College of Chest Physicians,Chicago Ill., October 1983.

[0014] Although CPAP has been found to be very effective and wellaccepted, it suffers from some of the same limitations, although to alesser degree, as do the surgical options; specifically, a significanproportion of sleep apnea patients do not tolerate CPAP well. Thus,development of other viable non-invasive therapies has been a continuingobjective in the art.

BRIEF SUMMARY OF THE INVENTION

[0015] The present invention contemplates a novel and improved methodfor treatment of sleep apnea as well as novel methodology and apparatusfor carrying out such improved treatment method. The inventioncontemplates the treatment of sleep apnea through application ofpressure at variance with ambient atmospheric pressure within the upperairway of the patient in a manner to promote dilation of the airway tothereby relieve upper airway occlusion during sleep.

[0016] In one embodiment of the invention, positive pressure is appliedalternately at relatively higher and lower pressure levels within theairway of the patient so that the pressure-induced force applied todilate the patient's airway is alternately a larger and a smallermagnitude dilating force. The higher and lower magnitude positivepressures are initiated by spontaneous patient respiration with thehigher magnitude pressure being applied during inspiration and the lowermagnitude pressure being applied during expiration.

[0017] The invention further contemplates a novel and improved apparatuswhich is operable in accordance with a novel and improved method toprovide sleep apnea treatment. More specifically, a flow generator andan adjustable pressure controller supply air flow at a predetermined,adjustable pressure to the airway of a patient through a flowtransducer. The flow transducer generates an output signal which is thenconditioned to provide a signal proportional to the instantaneous flowrate of air to the patient. The instantaneous flow rate signal is fed toallow pass filter which passes only a signal indicative of the averageflow rate over time. The average flow rate signal typically would beexpected to be a value representing a positive flow as the system islikely to have at least minimal leakage from the patient circuit (e.g.small leaks about the perimeter of the respiration mask worn by thepatient). The average flow signal is indicative of leakage because thesummation of all other components of flow over time must be essentiallyzero since inspiration flow must equal expiration flow volume over time,that is, over a period of time the volume of air breathed in equals thevolume of the gases breathed out.

[0018] Both the instantaneous flow signal and the average flow ratesignal are fed to an inspiration/expiration decision module which is, inits simplest form, a comparator that continually compares the inputsignals and provides a corresponding drive signal to the pressurecontroller. In general, when the instantaneous flow exceeds averageflow, the patient is inhaling and the drive signal supplied to thepressure controller sets the pressure controller to deliver air, at apreselected elevated pressure, to the airway of the patient. Similarly,when the instantaneous flow rate is less than the average flow rate, thepatient is exhaling-and the decision circuitry thus provides a drivesignal-to set the pressure controller to provide a relatively lowermagnitude of pressure in the airway of the patient. The patient's airwaythus is maintained open by alternating higher and lower magnitudes ofpressure which are applied during spontaneous inhalation and exhalation,respectively.

[0019] As has been noted, some sleep apnea patients do not toleratestandard CPAP therapy. Specifically, approximately 25% of patientscannot tolerate CPAP due to the attendant discomfort. CPAP mandatesequal pressures during both inhalation and exhalation. The elevatedpressure during both phases of breathing may create difficulty inexhaling and the sensation of an inflated chest. However, we havedetermined that although both inspiratory and expiratory air flowresistances in the airway are elevated during sleep preceding the onsetof apnea, the airway flow resistance may be less during expiration thanduring inspiration. Thus it follows that the bi-level CPAP therapy ofour invention as characterized above may be sufficient to maintainpharyngeal patency during expiration even though the pressure appliedduring expiration is not as high as that needed to maintain pharyngealpatency during inspiration. In addition, some patients may haveincreased upper airway resistance primarily during inspiration withresulting adverse, physiologic consequences. Thus, our invention alsocontemplates applying elevated pressure only during inhalation thuseliminating the need for global (inhalation and exhalation) increases inairway pressure. The relatively lower pressure applied during expirationmay in some cases approach or equal ambient pressure. The lower pressureapplied in the airway during expiration enhances patient tolerance byalleviating some of the uncomfortable sensations normally associatedwith CPAP.

[0020] Under prior CPAP therapy, pressures as high as 20 cm H2O havebeen required, and some patients on nasal CPAP thus have been needlesslyexposed to unnecessarily high expiratory pressures with the attendantdiscomfort and elevated mean airway pressure, and theoretic risk ofbarotrauma. Our invention permits independent application of a higherinspiratory airway pressure in conjunction with a lower expiratoryairway pressure in order to provide a therapy which is better toleratedby the 25% of the patient population which does not tolerat CPAPtherapy, and which may be safer in the other 75% of the patientpopulation.

[0021] As has been noted hereinabove, the switch between higher andlower pressure magnitudes can be controlled by spontaneous patientrespiration, and the patient thus is able to independently governrespiration rate and volume. As has been also noted, the inventioncontemplates automatic compensation for system leakage whereby nasalmask fit and air flow system integrity are of less consequence than inthe prior art. In addition to the-benefit of automatic leakcompensation, other important benefits of the invention include lowermean airway pressures for the patient and enhanced safety, comfort andtolerance.

[0022] It is accordingly one object of the present invention to providenovel and improved method for treatment of sleep apnea.

[0023] A further object of the invention to provide a novel and improvedapparatus which is operable according to a novel methodology in carryingout such a treatment for sleep apnea.

[0024] Another object of the invention is to provide a method andapparatus for treating sleep apnea by application of alternately highand low magnitudes of pressure in the airway of the patient with thehigh and low pressure magnitudes being initiated by spontaneous patientrespiration.

[0025] Another object of the invention is to provide an apparatus forgenerating alternately high and low pressure gas flow to a consumer ofthe gas with the higher and lower pressure flows being controlled bycomparison of the instantaneous flow rate to the gas consumer with theaverage flow rate to the consumer, which average flow rate may includeleakage from the system, and whereby the apparatus automaticallycompensates for system leakage.

[0026] These and other objects and further advantages of the inventionwill be more readily appreciated upon consideration of the followingdetailed description and accompanying drawings, in which:

[0027]FIG. 1 is a functional block diagram of an apparatus according tothe instant invention which is operable according to the method of theinstant invention;

[0028]FIG. 2 is a functional block diagram showing an alternativeembodiment of the invention;

[0029]FIG. 3 is a functional block diagram of the Estimated LeakComputer of FIG. 2;

[0030]FIG. 4 is a frontal elevation of a control panel for the apparatusof this invention;

[0031]FIG. 5 is a trace of patient flow rate versus time pertaining toanother embodiment of the invention;

[0032]FIG. 6 is a flow diagram relating to the embodiment of FIG. 5;

[0033]FIG. 7 is a trace of patient flow rate versus time illustratinganother embodiment of the invention;

[0034]FIG. 8 is a schematic illustration of a circuit according to anembodiment of the invention;

[0035]FIG. 9 is a trace of flow rate versus time showing processed andunprocessed flow rate signals;

[0036]FIGS. 10a and 10 b are illustrations of the operation of controlalgorithms according to other embodiments of the invention;

[0037]FIGS. 11a and 11 b are illustrations of the operation of stillother control algorithms according to other embodiments of theinvention; and

[0038]FIG. 12 illustrates operation of yet another control algorithmaccording to the invention.

[0039] There is generally indicated at 10 in FIG. 1 an apparatusaccording to one presently preferred embodiment of the instant inventionand shown in the form of a functional block diagram. Apparatus 10 isoperable according to a novel process which is another aspect of theinstant invention for delivering breathing gas such as air alternatelyat relatively higher and lower pressures (i.e., equal to or aboveambient atmospheric pressure) to a patient 12 for treatment of thecondition known as sleep apnea.

[0040] Apparatus 10 comprises a gas flow generator 14 (e.g., a blower)which receives breathing gas from any suitable source, a pressurizedbottle 16 or the ambient atmosphere, for example. The gas flow from flowgenerator 14 is passed via a delivery conduit 20 to a breathingappliance such as a mask 22 of any suitable known construction which isworn by patient 12. The mask 22 may preferably be a nasal mask or a fullface mask 22 as shown. Other breathing appliances which may be used inlieu of a mask include nasal cannulae, an endotracheal tube, or anyother suitable appliance for interfacing between a source of breathinggas and a patient, consistent with the desired effect to be achievedthrough use of the apparatus 10.

[0041] The mask 22 includes a suitable exhaust port means, schematicallyindicated at 24, for exhaust of breathing gases during expiration.Exhaust port 24 preferably is a continuously open port which imposessuitable flow resistance upon exhaust gas flow to permit a pressurecontroller 26, located in line with conduit 20 between flow generator 14and mask 22, to control the pressure of air flow within conduit 20 andthus within the airway of the patient 12. For example, exhaust port 24may be of sufficient cross-sectional flow area to sustain a continuousexhaust flow of approximately 15 liters per minute. The flow via exhaustport 24 is one component, and typically the major component of theoverall system leakage, which is an important parameter of systemoperation. In an alternative embodiment to be discussed hereinbelow, ithas been found that a non-rebreathing valve may be substituted for thecontinuously open port 24.

[0042] The pressure controller 26 is operative to control the pressureof breathing gas within the conduit 20 and thus within the airway of thepatient. Pressure controller 26 is located preferably, although notnecessarily, downstream of flow generator 14 and may take the form of anadjustable valve which provides a flow path which is open to the ambientatmosphere via a restricted opening, the valve being adjustable tomaintain a constant pressure drop across the opening for all flow ratesand thus a constant pressure within conduit 20.

[0043] Also interposed in line with conduit 20, preferably downstream ofpressure controller 26, is a suitable flow transducer 28 which generatesan output signal that is fed as indicated at 29 to a flow signalconditioning circuit 30 for derivation of a signal proportional to theinstantaneous flow rate of breathing gas within conduit 20 to thepatient.

[0044] It will be appreciated that flow generator 14 is not necessarilya positive displacement device. It may be, for example, a blower whichcreates a pressure head within conduit 20 and provides air flow only tothe extent required to maintain that pressure head in the presence ofpatient breathing cycles, the exhaust opening 24, and action of pressurecontroller 26 as above described. Accordingly, when the patient isexhaling, peak exhalation flow rates from the lungs may far exceed theflow capacity of exhaust port 24. As a result, exhalation gas backflowswithin conduit 20 through flow transducer 28 and toward pressurecontroller 26, and the instantaneous flow rate signal from transducer 28thus will vary widely within a range from relatively large positive(i.e.toward the patient) flow to relatively large negative (i.e. fromthe patient) flow.

[0045] The instantaneous flow rate signal from flow signal conditioningcircuitry 30 is fed as indicated at 32 to a decision module 34, a knowncomparator circuit for example, and is additionally fed as indicated at36 to a low pass filter 38. Low pass filter 38 has a cut-off frequencylow enough to remove from the instantaneous flow rate input signal mostvariations in the signal which are due to normal breathing. Low passfilter 38 also has a long enough time constant to ensure that spurioussignals, aberrant flow patterns and peak instantaneous flow rate valueswill not dramatically affect system average flow. That is, the timeconstant of low pass filter 38 is selected to be long enough that itresponds slowly to the instantaneous flow rate signal input.Accordingly, most instantaneous flow rate input signals which could havea-large impact on system average flow in the short term have a muchsmaller impact over a longer term, largely because such instantaneousflow rate signal components will tend to cancel over the longer term.For example, peak instantaneous flow rate values will tend to bealternating relatively large positive and negative flow valuescorresponding to peak inhalation and exhalation flow achieved by thepatient during normal spontaneous breathing. The output of low passfilter 38 thus is a signal which is proportional to the average flow inthe system, and this is typically a positive flow which corresponds toaverage system leakage (including flow from exhaust 24) since, as noted,inhalation and exhalation flow cancel for all practical purposes.

[0046] The average flow signal output from the low pass filter 38 is fedas indicated at 40 to decision circuitry 34 where the instantaneous flowrate signal is continually compared to the system average flow signal.The output of the decision circuitry 34 is fed as a drive signalindicated at 42 to control the pressure controller 26. The pressuremagnitude of breathing gas within conduit 20 thus is coordinated withthe spontaneous breathing effort of the patient 12, as follows.

[0047] When the patient begins to inhale, the instantaneous flow ratesignal goes to a positive value above the positive average flow signalvalue. Detection of this increase in decision circuitry 34 is sensed asthe start of patient inhalation. The output signal from decisioncircuitry 34 is fed to pressure controller 26 which, in response,provides higher pressure gas flow within conduit 20 and thus higherpressure within the airway of the patient 12. This is the highermagnitude pressure value of our bi-level CPAP system and is referred tohereinbelow as IPAP (inhalation positive airway pressure). Duringinhalation, the flow rate within conduit 20 will increase to a maximumand then decrease as inhalation comes to an end.

[0048] At the start of exhalation, air flow into the patient's lungs isnil and as a result the instantaneous flow rate signal will be less thanthe average flow rate signal which, as noted is a relatively constantpositive flow value. The decision circuitry 34 senses this condition asthe start of exhalation and provides a drive signal to pressurecontroller 26 which, in response, provides gas flow within conduit 20 ata lower pressure which is the lower magnitude pressure value of thebi-level CPAP system, referred to hereinbelow as EPAP (exhalationpositive airway pressure). As has been noted hereinabove the range ofEPAP pressures may include ambient atmospheric pressure. When thepatient again begins spontaneous inhalation, the instantaneous flow ratesignal again increases over the average flow rate signal, and thedecision circuitry once again feeds a drive signal to pressurecontroller 26 to reinstitute the IPAP pressure.

[0049] System operation as above specified requires at least periodiccomparison of the input signals 32 and 40 by decision circuitry 34.Where this or other operations are described herein as continual, thescope of meaning to be ascribed includes both continuous (i.e.uninterrupted) and periodic (i.e. at discrete intervals).

[0050] As has been noted, the system 10 has a built-in controlledleakage via exhaust port 24 thus assuring that the average flow signalnormally will be at least a small positive flow, although in somecircumstances such as when oxygen is added to the gas flow, the averageflow may be a negative value. During inhalation, the flow sensed by theflow transducer will be the sum of exhaust flow via port 24 and allother system leakage downstream of transducer 28, and inhalation flowwithin the airway of the patient 12. Accordingly, during inhalation theinstantaneous flow rate signal as conditioned by conditioning module 30,will reliably and consistently reflect inhalation flow exceeding theaverage flow rate signal. During exhalation, the flow within conduit 20reverses as exhalation flow from the lungs of the patient far exceedsthe flow capacity of exhaust port 24. Accordingly, exhalation airbackflows within conduit 20 past transducer 28 and toward pressurecontroller 26. Since pressure controller 26 is operable to maintain setpressure, it will act in response to flow coming from both the patientand the flow generator to open an outlet port sufficiently toaccommodate the additional flow and thereby maintain the specified setpressure as determined by action of decision circuitry 34.

[0051] In both the inhalation and exhalation cycle phases, the pressureof the gas within conduit 20 exerts a pressure within the airway of thepatient to maintain an open airway and thereby alleviate airwayconstriction.

[0052] In practice, it may be desirable to provide a slight offset inthe switching level within decision circuitry 34 with respect to theaverage flow rate signal, so that the system does not prematurely switchfrom the low pressure exhalation mode to the higher pressure inhalationmode. That is, a switching setpoint offset in the positive directionfrom system average flow may be provided such that the system will notswitch to the IPAP mode until the patient actually exerts a significantspontaneous inspiratory effort of a minimum predetermined magnitude.This will ensure that the initiation of inhalation is completelyspontaneous and not forced by an artificial increase in airway pressure.A similar switching setpoint offset may be provided when in the IPAPmode to ensure the transition to the lower pressure EPAP mode will occurbefore the flow rate of air into the lungs of the patient reaches zero(i.e. the switch to EPAP occurs slightly before the patient ceasesinhalation.) This will ensure that the patient will encounter no undueinitial resistance to spontaneous exhalation.

[0053] From the above description, it will be seen that a novel methodof treating sleep apnea is proposed according to which the airwaypressure of the patient is maintained at a higher positive pressureduring inspiration and a relatively lower pressure during expiration,all without interference with the spontaneous breathing of the patient.The described apparatus is operable to provide such treatment for sleepapnea patients by providing a flow of breathing gas to the patient atpositive pressure, and varying the pressure of the air flow to providealternately high and low pressure within the airway of the patientcoordinated with the patient's spontaneous inhalation and exhalation.The described system can also be used wit pressure support systems suchas the proportional airway pressure system described in U.S. Pat. No.5,107,830 of Younes, the entire disclosure of which is herebyincorporated herein and made a part hereof by reference.

[0054] To provide pressure control, the flow rate of breathing gas tothe patient is detected and processed to continually provide a signalwhich is proportional to the instantaneous breathing gas flow rate inthe system. The instantaneous flow rate signal is further processed toeliminate variations attributable to normal patient respiration andother causes thus generating a signal which is proportional to theaverage or steady state system gas flow. The average flow signal iscontinually compared with the instantaneous flow signal as a means todetect the state of the patient's spontaneous breathing versus averagesystem flow. When instantaneous flow exceeds the average flow, thepatient is inhaling, and in response the pressure of gas flowing to thepatient is set at a selected positive pressure, to provide acorresponding positive pressure within the airway of the patient. Whencomparison of the instantaneous flow rate signal with the average flowsignal indicates the patient is exhaling, as for example when theinstantaneous flow signal indicates flow equal to or less than theaverage flow, the pressure of breathing gas to the patient is adjustedto a selected lower pressure to provide a corresponding lower pressurewithin the airway of the patient.

[0055] In an alternative embodiment of the invention as shown in FIGS. 2and 3, the low pass filter 38 is replaced by an estimated leak computerwhich includes a low pass filter as well as other functional elements asshown in FIG. 3. The remainder of the system as shown in FIG. 2 issimilar in most respects to the system shown in FIG. 1. Accordingly,like elements are identified by like numbers, and the descriptionhereinabove of FIG. 1 embodiment also applies generally to FIG. 2.

[0056] By using the operative capability of the estimated leak computer50, as described hereinbelow, it is possible to adjust the referencesignal which is fed to decision circuitry 34 on a breath by breath basisrather than merely relying on long term average system flow. Todistinguish this new reference signal from average system flow it willbe referred to hereinbelow as the estimated leak flow rate signal orjust the estimated leak signal.

[0057] As was noted hereinabove, the average system flow rate referencesignal changes very slowly due to the long time constant of the low passfilter 38. This operative feature was intentionally incorporated toavoid disturbance of the reference signal by aberrant instantaneous flowrate signal inputs such as erratic breathing patterns. While it waspossible to minimize the impact of such aberrations on the average flowrate reference signal, the average flow signal did nevertheless change,although by small increments and only very slowly in response todisturbances. Due to the long time constant of the low pass filter, suchchanges in the reference signal even if transitory could last for a longtime.

[0058] Additionally, even a small change in the reference signal couldproduce a very significant effect on system triggering. For example,since the objective is to trigger the system to the IPAP mode wheninhalation flow just begins to go positive, small changes in thereference signal could result in relatively large changes in thebreathing effort needed to trigger the system to the IPAP mode. In someinstances the change in reference signal could be so great that withnormal breathing effort the patient would be unable to trigger thesystem. For example, if the system were turned on before placement ofthe mask on the face of the patient, the initial free flow of air fromthe unattached mask could result in a very large magnitude positivevalue for initial average system flow. If such value were to exceed themaximum inspiratory flow rate achieved in spontaneous respiration by thepatient, the system would never trigger between the IPAP and EPAP modesbecause the decision circuitry would never see an instantaneous flowrate signal greater than the average flow rate signal, at least notuntil a sufficient number of normal breathing cycles after applicationof the mask to the patient to bring the reference signal down to a valuemore closely commensurate with the actual system leak in operation. Ashas been noted, with the low pass filter this could take a rather longtime, during which time the patient would be breathing spontaneouslyagainst a uniform positive pressure. This would not be at all in keepingwith the present invention.

[0059] In addition to the embodiment based on a reference signal derivedfrom estimated leak flow rate on a breath by breath basis which iscontrolled totally by spontaneous patient breathing, two further mode ofoperation also are envisioned, one being spontaneous/timed operation inwhich the system automatically triggers to the IPAP mode for just longenough to initiate patient inspiration if the system does not senseinspiratory effort within a selected time after exhalation begins. Toaccomplish this,.a timer is provided which is reset at the beginning ofeach patient inspiration whether the inspiratory cycle was triggeredspontaneously or by the timer itself. Thus, only the start ofinspiration is initiated by the timer. The rest of the operating cyclein this mode is controlled by spontaneous patient breathing and thecircuitry of the system to be described.

[0060] A further mode of operation is based purely on timed operation ofthe system rather than on spontaneous patient breathing effort, but withtimed cycles in place of spontaneous patient breathing.

[0061] Referring to FIG. 3, the estimated leak computer 50 includes thelow pass filter 38′ as well as other circuits which are operative tomake corrections to the estimated leak flow rate signal based onon-going analysis of each patient breath. A further circuit is providedwhich is operative to adjust the estimated leak flow rate signal quicklyafter major changes in system flow such as when the blower has beenrunning prior to the time when the mask is first put on the patient, orafter a major leak in the system has either started or has been shutoff.

[0062] The low pass filter 38′ also includes a data storage capabilitywhose function will be described hereinbelow.

[0063] The low pass filter 38′ operates substantially as described abovewith reference to FIG. 1 in that it provides a long term average ofsystem flow which is commensurate with steady state system leakageincluding the flow capacity of the exhaust port 24. This long termaverage is operative in the FIG. 3 embodiment to adjust the estimatedleak flow rate reference signal only when system flow conditions arechanging very slowly.

[0064] To provide breath by breath analysis and adjustment of thereference signal, a differential amplifier 52 receives the instantaneousflow rate signal as indicated at 54, and the estimated leak signaloutput from low pass filter 38′ as indicated at 56.

[0065] The output of differential amplifier 52 is the difference betweeninstantaneous flow rate and estimated leak flow rate, or in other wordsestimated instantaneous patient flow rate. This will be clear uponconsidering that instantaneous flow is the su of patient flow plusactual system leakage. The estimated patient flow signal output fromdifferential amplifier 52 is provided as indicated at 58 to a flowintegrator 60 which integrates estimated patient flow breath by breathbeginning and ending with the trigger to IPAP. Accordingly, anadditional input to the flow integrator 60 is the IPAP/EPAP state signalas indicated at 62. The IPAP/EPAP state signal is the same as the drivesignal provided to pressure controller 26; that is, it is a signalindicative of the pressure state, as between IPAP and EPAP, of thesystem. The state signal thus may be used to mark the beginning and endof each breath for purposes of breath by breath integration byintegrator 60.

[0066] If the estimated leak flow rate signal from low pass filter 38′is equal to the true system leak flow rate, and if the patient's inhaledand exhaled volumes are identical for a given breath (i.e. totalpositive patient flow equals total negative patient flow for a givenbreath), then the integral calculated by integrator 60 will be zero andno adjustment of estimated leak flow rate will result. When the integralcalculated by integrator 60 is non-zero, the integral value in the formof an output signal from integrator 60 is provided as indicated at 64 toa sample and hold module 66. Of course, even with a zero value integral,an output signal may be provided to module 66, but the ultimate resultwill be no adjustment of the estimated leak flow rate signal.

[0067] A non-zero integral value provided to module 66 is furtherprovided to module 38′ as indicated at 68 with each patient breath byoperative action of the IPAP/EPAP state signal upon module 66 asindicated at 70. The effect of a non-zero integral value provided tomodule 38′ is an adjustment of the estimated leak flow rate signalproportional to the integral value and in the direction which wouldreduce the integral value towards zero on the next breath if all otherconditions remain the same.

[0068] With this system, if the patient's net breathing cycle volume iszero, and if the system leak flow rate changes, the integrator circuitwill compensate for the change in leak flow rate by incrementaladjustments to the estimated leak flow rate within several patientbreaths.

[0069] The integrator circuit 60 also will adjust the estimated leakflow rate signal in response to non-zero net volume in a patientbreathing cycle. It is not unusual for a patient's breathing volume tobe non-zero. For example, a patient may inhale slightly more on eachbreath than he exhales over several breathing cycles, and then followwith a deeper or fuller exhalation. In this case, the integrator circuitwould adjust the estimated leak flow rate signal as if the actual systemleak rate had changed; however, since the reference signal correction isonly about one tenth as large as would be required to make the totalcorrection in one breath, the reference signal will not changeappreciably over just one or two breaths. Thus, the integrator circuitaccommodates both changes in system leakage and normal variations inpatient breathing patterns.

[0070] An end exhalation module 74 is operative to calculate anotherdata component for use in estimating the system leak flow rate asfollows. The module 74 monitors the slope of the instantaneous flow ratewave form. When the slope value is near zero during exhalation (asindicated by the state signal input 76) the indication is that the flowrate is not changing. If the slope of the instantaneous flow rate signalwave form remains small after more than one second into the respiratoryphase, the indication is that exhalation has ended and that the net flowrate at this point thus is the leak flow rate. However, if estimatedpatient flow rate is non-zero at the same time, one component of theinstantaneous flow rate signal must be patient flow.

[0071] When these conditions are met, the circuit adjusts the estimatedleak flow rate slowly in a direction to move estimated patient flow ratetoward zero to conform to instantaneous patient flow conditions expectedat the end of exhalation. The adjustment to estimated leak flow rate isprovided as an output from module 74 to low pass filter 38′ as indicatedat 80. When this control mechanism takes effect, it disables the breathby breath volume correction capability of integrator circuit 60 for thatbreath only.

[0072] The output of module 74 is a time constant control signal whichis provided to low pass filter 38′ to temporarily shorten the timeconstant thereof for a sufficient period to allow the estimated leakflow rate to approach the instantaneous flow rate signal at thatspecific instant. It will be noted that shortening the low pass filtertime constant increases the rapidity with which the low pass filteroutput (a system average) can adjust toward the instantaneous flow ratesignal input.

[0073] Another component of estimated leak flow rate control is a grosserror detector 82 which acts when the estimated patient flow rate,provided thereto as indicated at 84, is away from zero for more thanabout 5 seconds. Such a condition may normally occur, for example, whenthe flow generator 14 is running before mask 22 is applied to thepatient. This part of the control system is operative to stabilizeoperation quickly after major changes in the leak rate occur.

[0074] In accordance with the above description, it will be seen thatlow pass filter 38′ acts on the instantaneous flow rate signal toprovide an output corresponding to average system flow, which is systemleakage since patient inspiration and expiration over time constitutes anet positive flow of zero. With other enhancements, as described, thesystem average flow can be viewed as an estimate of leakage flow rate.

[0075] The differential amplifier 52 processes the instantaneous flowrate signal and the estimated leak flow rate signal to provide anestimated patient flow rate signal which is integrated and non-zerovalues of the integral are fed back to module 38′ to adjust theestimated leak flow rate signal on a breath by breath basis. Theintegrator 60 is reset by the IPAP/EPAP state signal via connection 62.

[0076] Two circuits are provided which can override the integratorcircuit, including end exhalation detector 74 which provides an outputto adjust the time constant of low pass filter 38′ and which also isprovided as indicated at 86 to reset integrator 60. Gross error detector82 is also provided to process estimated patient flow rate and toprovide an adjustment to estimated leak flow rate under conditions asspecified. The output of module 82 also is utilized as an integratorreset signal as indicated at 86. It will be noted that the integrator 60is reset with each breath of the patient if, during that breath, it isultimately overridden by module 74 or 82. Accordingly, the multiplereset capabilities for integrator 60 as described are required.

[0077] In operation, the system may be utilized in a spontaneoustriggering mode, a spontaneous/timed mode or a purely timed mode ofoperation. In spontaneous operation, decision circuitry 34 continuouslycompares the instantaneous flow rate with estimated leak flow rate. Ifthe system is in the EPAP state or mode, it remains there untilinstantaneous flow rate exceeds estimated leak flow rate byapproximately 40 cc per second. When this transition occurs, decisioncircuitry 34 triggers the system into the IPAP mode for 150milliseconds. The system will then normally remain in the IPAP mode asthe instantaneous flow rate to the patient will continue to increaseduring inhalation due to spontaneous patient effort and the assistanceof the increased IPAP pressure.

[0078] After the transition to the IPAP mode in each breath, a temporaryoffset is added to the estimated leak flow rate reference signal. Theoffset is proportional to the integral of estimated patient flow ratebeginning at initiation of the inspiratory breath so that it graduallyincreases with time during inspiration at a rate proportional to thepatient's inspiratory flow rate. Accordingly, the flow rate level aboveestimated leak flow needed to keep the system in the IPAP mode duringinhalation decreases with time from the beginning of inhalation and inproportion to the inspiratory flow rate, With this enhancement, thelonger an inhalation cycle continues, the larger is the reference signalbelow which instantaneous flow would have to decrease in order totrigger the EPAP mode. For example, if a patient inhales at a constant500 cc per second until near the end of inspiration, a transition toEPAP will occur when his flow rate drops to about 167 cc per secondafter one second, or 333 cc per second after two seconds, or 500 cc persecond after three seconds, and so forth. For a patient inhaling at aconstant 250 cc per second, the triggers would occur at 83, 167 and 250cc per second at one, two and three seconds into IPAP, respectively.

[0079] In this way, the EPAP trigger threshold comes up to meet theinspiratory flow rate with the following benefits. First, it becomeseasier and easier to end the inspiration cycle with increasing time intothe cycle. Second, if a leak develops which causes an increase ininstantaneous flow sufficient to trigger the system into the IPAP mode,this system will automatically trigger back to the EPAP mode after about3.0 seconds regardless of patient breathing effort. This would allow thevolume-based leak correction circuit (i.e. integrator 60) to act as itis activated with each transition to the IPAP mode. Thus, if a leakdevelops suddenly, there will be a tendency toward automatic triggeringrather than spontaneous operation for a few breaths, but the circuitwill not be locked into the IPAP mode.

[0080] Upon switching back to the EPAP mode, the trigger threshold willremain above the estimated leak flow rate for approximately 500milliseconds to allow the system to remain stable in the EPAP modewithout switching again -while the respective flow rates are changing.After 500 milliseconds, the trigger threshold offset is reset to zero toawait the next inspiratory effort.

[0081] The normal state for the circuit is for it to remain in the EPAPmode until an inspiratory effort is made by the patient. The automaticcorrections and adjustments to the reference signal are effective tokeep the system from locking up in the IPAP mode and to preventauto-triggering while at the same time providing a high level ofsensitivity to inspiratory effort and rapid adjustment for changing leakconditions and breathing patterns.

[0082] In the spontaneous/timed mode of operation, the system performsexactly as above described with reference to spontaneous operation,except that it allows selection of a minimum breathing rate to besuperimposed upon the spontaneous operating mode. If the patient doesnot make an inspiratory effort within a predetermined time, the systemwill automatically trigger to the IPAP mode for 200 milliseconds. Theincreased airway pressure for this 200 milliseconds will initiatepatient inspiration and provide sufficient time that spontaneous patientflow will exceed the reference signal so that the rest of the cycle maycontinue in the spontaneous mode as above described. The breaths perminute timer is reset by each trigger to IPAP whether the transition wastriggered by the patient or by the timer itself.

[0083] In the timed operating mode, all triggering between IPAP and EPAPmodes is controlled by a timer with a breaths per minute control beingused to select a desired breathing rate from, for example, 3 to 30breaths per minute. If feasible, the selected breathing rate iscoordinated to the patient's spontaneous breathing rate. The percentIPAP control is used to set the fraction of each breathing cycle to bespent in the IPAP mode. For example, if the breaths per minute controlis set to 10 breaths per minute (6 seconds per breath) and the percentIPAP control is set to 33%, then the flow generator will spend, in eachbreathing cycle, two seconds in IPAP and four seconds in EPAP.

[0084]FIG. 4 illustrates a control panel for controlling the systemabove described and including a function selector switch which includesfunction settings for the three operating modes of spontaneous,spontaneous/timed, and timed as above described. The controls forspontaneous mode operation include IPAP and EPAP pressure adjustmentcontrols 90 and 92, respectively. These are used for setting therespective IPAP and EPAP pressure levels. In the spontaneous/timed modeof operation, controls 90 and 92 are utilized as before to set IPAP andEPAP pressure levels, and breaths per minute control 94 additionally isused to set the minimum desired breathing rate in breaths per minute. Inthe timed mode of operation, controls 90, 92 and 94 are effective, andin addition the percent IPAP control 96 is used to set the timepercentage of each breath to be spent in the IPAP mode.

[0085] Lighted indicators such as LED's 96, 98 and 100 are also providedto indicate whether the system is in the IPAP or EPAP state, and toindicate whether in the spontaneous/timed mode of operation theinstantaneous state of the system is spontaneous operation or timedoperation.

[0086] An alternative embodiment of the invention contemplates detectingthe beginning of inspiration and expiration by reference to a patientflow rate wave form such as shown in FIG. 5. Comparison of instantaneousflow rate with average flow rate as set forth hereinabove provides asatisfactory method for determining whether a patient is inhaling orexhaling; however, other means for evaluating instantaneous flow ratecan also be used, and these may be used alone or in combination withaverage flow rate.

[0087] For example, FIG. 5 shows a typical patient flow rate wave formwith inspiration or inhalation flow shown as positive flow above baseline B and exhalation flow shown as negative flow below base line B. Theflow rate wave form may thus be sampled at discrete time intervals. Thecurrent sample is compared with that taken at an earlier time. Thisapproach appears to offer the benefit of higher sensitivity to patientbreathing effort in that it exhibits less sensitivity to errors in theestimated system leakage, as discussed hereinabove.

[0088] Normally, the estimated inspiratory and expiratory flow rate waveforms will change slowly during the period beginning after a few hundredmilliseconds into the respective inspiratory and expiratory phases, upuntil the respective phase is about to end. Samples of the flow ratewave form are taken periodically, and the current sample is comparedrepeatedly with a previous sample. During inspiration, if the magnitudeof the current sample is less than some appropriate fraction of thecomparison sample, then the inspiration phase is deemed to be finished.This condition thus can be used to trigger a change to the desiredexhalation phase pressure. The same process can be used duringexhalation to provide a trigger condition for the changeover toinhalation pressure.

[0089]FIG. 6 illustrates a flow diagram for a suitable algorithm showingrepeated sampling of the patient flow rate wave form at time intervalsΔt, as shown at 110, and repeated comparison of the flow at current timet with a prior sampling, for example the flow at time t-4 as indicatedat 112. Of course, the time increment between successive samplings issmall enough that the comparison with the value observed in the fourthprior sampling event covers a suitably small portion of the flow ratewave form. For example, as shown in FIG. 5 the designations for flowsampled at time t, the sample value taken four sampling events prior attime t-4, and the time interval Δt illustrate one suitable proportionaterelationship between Δt and the time duration of a patient breath. Thisis, however, only an illustration. In fact, electronic technology suchas disclosed hereinabove would be capable of performing this algorithmusing much smaller Δt intervals, provided other parameters of thealgorithm, such as the identity of the comparison sample and/or theproportion relationship between the compared samples, are adjustedaccordingly.

[0090] Continuing through the flow chart of FIG. 6, if the patient'sstate of breathing is inspiration (114), the current flow sample valueis compared with the sample value observed t-4 sampling events prior. Ifthe current flow sample value is less than, for example, 75% of thecomparison flow value (116), the system triggers a change in state tothe exhalation phase (118). If the current flow sample value is not lessthan 75% of the comparison value (120), no change in breathing state istriggered.

[0091] If the breathing state is not inspiration, and the current flowsample value is less than 75% of the selected comparison value (122),then an expiration phase, rather than an inspiration phase, is completedand the system is triggered to change the state of breathing toinspiration (124). If the specified flow condition is not met, again nochange in breathing phase is triggered (126).

[0092] The routine characterized by FIG. 6 repeats continuously asindicated at 128, each time comparing a current flow sample with thefourth (or other suitable) prior sample to determine whether the systemshould trigger from one breathing phase to the other.

[0093] Of course, it will be understood that the embodiment set forthabove with reference to FIGS. 5 and 6 does not trigger changes inbreathing per se within the context of the invention as set forth in thebalance of the above specification. Rather, in that context the FIGS. 5and 6 embodiment merely detects changes between inspiration andexpiration in spontaneous patient breathing, and triggers the system tosupply the specified higher or lower airway pressure as set forthhereinabove.

[0094] Common problems resulting from insufficient ventilator triggeringsensitivity include failure of the ventilator to trigger to IPAP uponexertion of inspiratory effort by the patient, and failure of theventilator in IPAP to trigger off or to EPAP at the end of patientinspiratory effort. Patient respiratory effort relates primarily to theinspiratory portion of the respiration cycle because in general only theinspiratory part of the cycle involves active effort. The expiratoryportion of the respiration cycle generally is passive. A furthercharacteristic of the respiratory cycle is that expiratory flow rategenerally may reach zero and remain at zero momentarily before the nextinspiratory phase begins. Therefore the normal stable state for apatient-triggered ventilator should be the expiratory state, which isreferred to herein as EPAP.

[0095] Although the apparatus as above described provides sensitivetriggering between the IPAP and EPAP modes, triggering can be stillfurther improved. The primary objectives of improvements in triggeringsensitivity generally are to decrease the tendency of a system totrigger automatically from EPAP to IPAP in the absence of inspiratoryeffort by the patient, and to increase system sensitivity to the end ofpatient inspiratory effort. These are not contradictor objectives.Decreasing auto-trigger sensitivity does not necessarily also decreasetriggering sensitivity to patient effort.

[0096] Generally, for a patient whose ventilatory drive is functioningnormally, the ideal ventilator triggering sensitivity is represented byclose synchronization between changes of state in the patient'srespiration (between inspiration and expiration), and the correspondingchanges of state by the ventilator. Many conventional pressure supportventilators which trigger on the basis of specified flow or pressurethresholds can approach such close synchronization only under limitedconditions.

[0097] In the system described hereinabove, where triggering is based onspecified levels of estimated patient flow rate, close synchronizationbetween system state changes and patient respiration state changes canbe readily achieved. This is especially true for patients having nosevere expiratory flow limitation as the flow rate reversals for suchpatients typically may correspond quite closely with changes in thepatient respiratory effort; however, flow rate reversals may notnecessarily correspond precisely to changes in patient respiratoryeffort for all patients. For example, those patients experiencingrespiratory distress or those being ventilated with pressure supportventilation may not exhibit the desired close correspondence betweenpatient effort and breathing gas supply flow rate reversals.

[0098] In addressing improvements in synchronized triggering based onflow rate, at least three different respiratory cycles are consideredThe first is the patient respiratory cycle as indicated by patienteffort. If the patient makes any effort, it must include inspiratoryeffort. Patient expiration may be active or passive.

[0099] The second cycle is the ventilator respiratory cycle, that is,the cycle delivered by the ventilator. This cycle also has aninspiratory phase and an expiratory-phase, but the ventilator may or maynot be synchronized with the corresponding phases of the patientrespiratory cycle.

[0100] The third cycle is the flow respiratory cycle, as indicated bythe direction of flow to or from the patient's airway. Flow passes fromthe ventilator into the patient on inspiration and out of the patient onexpiration. If one looks at this cycle only, the inspiratory/expiratorychanges could be identified by the zero flow crossings between positiveand negative flow.

[0101] Ideally, if the patient's breathing drive is competent, the flowrespiratory cycle should be synchronized with the patient respiratorycycle, with the assistance of the ventilator. One important objective ofthe triggering described here is to synchronize the ventilator cyclewith the patient cycle so that the flow cycle is coordinated withpatient inspiratory effort. With flow triggering the pressure deliveredby the ventilator affects the flow, in addition to the effect on flow ofthe patient's own efforts. This is another reason why zero flowcrossings may not necessarily be good indicators of changes in patienteffort.

[0102] As described, flow triggering provides certain advantages notavailable with some other modes of ventilator triggering. For example,in ventilators which use pressure-variation for triggering, when thepatient makes an inspiratory effort lung volume begins to increase. Thisvolume change causes a pressure drop, which in some conventional closedcircuit ventilators is sensed by the ventilator triggering system toinitiate breathing gas delivery. With flow triggering there is no needfor a pressure drop to occur in the breathing circuit. The triggeringalgorithm may still require the patient lung volume to increase, butsince the pressure in the ventilator circuit does not have to drop thepatient expends less energy in the inspiratory effort to achieve a givenvolume change.

[0103] Additionally, in an open circuit, that is one in which exhaledgases are flushed from the system via a fixed leak, it would bedifficult to generate the required pressure drop for pressure basedtriggering upon exertion of small patient inspiratory effort. That is,with an open circuit the magnitude of patient effort required togenerate even a very small pressure drop would be considerable. Thus, itwould be correspondingly quite difficult to achieve sufficientsensitivity in a pressure based triggering system to provide reliabletriggering if used in conjunction with an open circuit. In general, theopen circuit is simpler than one with a separate exhaust valve andexhaust tubing and therefore simpler to use and less costly, and forthese and other reasons may often be preferred over a closed circuit.

[0104] In a presently preferred triggering scheme for the abovedescribed flow rate triggering system, the flow rate signal must exceeda threshold value equivalent to 40 cc per second continuously for 35milliseconds in order to trigger the system to IPAP. If the flow ratesignal drops below the specified threshold value during the specifiedtime interval, the threshold timer is reset to zero. These thresholdvalues permit triggering with a minimal patient volume change of only1.4 cc, which represents much improved sensitivity over some priorpressure triggered ventilators.

[0105] In reality, however, such minimal patient volume changerepresents overly sensitive triggering which would not be practical inactual practice. Noise in the flow rate signal would continually resetthe threshold timer during intervals when the patient flow rate is veryclose to 40 cc per second. Therefore, it is believed a minimum volumechange for consistent triggering of the described system would be about3 cc., although this too is considered to be a very sensitive triggeringlevel. It is noted, however, that pressure signal noise can also occurin prior pressure triggered ventilators, thus creating triggeringsensitivity problems. Due to this and other details of their operationthe theoretical maximum sensitivity of such prior ventilators is betterthan the sensitivity which can actually be achieved in operation.

[0106] As one example of an operating detail which can adversely affectsensitivity in prior ventilators, if the ventilator response time isslow, the resulting time delay between the time when a triggering timerequirement is satisfied and the time when gas supply pressure actuallybegins to rise will adversely affect sensitivity. During the time delay,the patient is continuing to make inspiratory effort and thus will bedoing much greater actual work than that needed to trigger theventilator, and much greater work than would be done if the ventilatorresponded more quickly. Another example occurs during rapid breathingwhen a patient begins to make an inspiratory effort before exhalationflow rate drops to zero. In this case, an additional delay occurs untilflow reverses and the exhalation valve has time to close.

[0107] Triggering to EPAP upon exhalation also can present problems. Atypical prior pressure support ventilator will trigger off at the end ofinspiration based on the patient's flow decreasing past a flowthreshold, or on timing out of a timer, for example. Thetrigger-to-exhalation flow threshold can be either a fixed value or afraction of peak flow, for example 25% of peak flow. As one example ofthe sort of problem which can affect triggering to exhalation in priorventilators, a patient with moderately severe COPD (chronic obstructivepulmonary disease) will exhibit comparatively high respiratory systemresistance and perhaps low respiratory system elastance. This willresult in a comparatively long time constant for the patient, the timeconstant being the time taken for exhalation flow to decay to ¼ of peakflow or other suitable selected proportion of peak flow. The long timeconstant for the COPD patient means that flow will drop slowly duringinspiration.

[0108] The only way this patient will be able to trigger the ventilatorto exhalation will be to increase exhalatory effort. That is, thepatient must actively decrease the flow rate by respiratory muscleeffort to trigger to exhalation. Further, because of the patient's largetime constant, there may be considerable exhalation flow when thepatient's expiratory cycle begins, and the trigger to IPAP thus will bedelayed, to the detriment of ventilator sensitivity.

[0109] To achieve greater sensitivity to patient effort in an improvedembodiment of the invention, the disclosed system may be provided withapparatus to continually monitor multiple conditions rather thanmonitoring only a single condition as represented by the abovedescription pertinent to element 82. In the improved embodiment,inspiratory time is limited to no more than about 3 seconds under normalconditions, and normal exhalation flow rate is required to be nearlyzero by about 5 seconds after the beginning of exhalation.

[0110] To employ the first of these conditions the system monitorsestimated patient flow rate and requires its value to cross zero atleast once approximately every 5 seconds. To monitor this condition, aninput is required from the output of a comparator which compares patientflow rate with a zero value. The comparator is a bistable device whichprovides one output (i.e. true) when patient flow is greater than zero,and an alternative output (i.e. false) when patient flow is less thanzero. If the comparator output is either true or false continuously forlonger than about 5 seconds, a signal is generated which triggers thesystem to EPAP.

[0111] The second condition is the monitoring of the IPAP/EPAP state tosee if it remains in EPAP for more than about 5 seconds. If it does, thevolume will be held at zero until there is another valid trigger toIPAP. This function prevents the volume signal from drifting away fromzero when there is an apnea. After 5 seconds in EPAP volume shouldnormally be zero.

[0112] Further improvements in triggering sensitivity requireconsideration of more than flow rate levels for determination ofsuitable triggering thresholds. In particular, changes in the shape orslope of the patient's flow rate wave form can be utilized to bettersynchronize a ventilation system with spontaneous respiratory effortexerted by the patient. Some such considerations are discussedhereinabove with reference to FIGS. 5 and 6.

[0113] A typical flow rate tracing for a normal patient experiencingrelaxed breathing is shown in FIG. 7. The simplest approach forsynchronizing ventilator triggering to spontaneous patient respirationhas often been to use the zero flow points 129 as the reference fortriggering between EPAP and IPAP; however, this approach does notaddress the potential problems. First, as noted above, flow ratereversals may not necessarily correspond to changes in patientspontaneous breathing effort for all patients. Second, the flow signalis not necessarily uniform. Concerning this problem in particular, smallflow variations corresponding to noise in electronic signals will occurdue to random variations in system and airway pressure caused by flowturbulence or random variations in the pressure control system.Additional flow rate noise results from small oscillations in thebreathing gas flow rate due to changes in blood volume within thepatient's chest cavity with each heart beat. Third, since pressuresupport commonly is an all-or-nothing mode of ventilation, and since therespiratory systems of all patients will exhibit an elastance component,an inappropriate trigger to IPAP will always result in delivery of somevolume of breathing gas unless the patient's airway is completelyobstructed. Thus, reliance on zero flow or flow reversals for triggeringwill in many instances result in loss of synchronization between apatient inspiratory effort and ventilator delivery of breathing gas.

[0114] Referring to FIG. 7, there is shown a flow rate versus time tracefor the respiratory flow of a typical, spontaneously breathing person.Time advances from left to right on the horizontal axis and flow ratevaries about a zero-value base line on the vertical axis. Of course, inthe description below referring to FIG. 7 it is to be understood thatthe illustrated flow rate versus time trace is an analog for spontaneousbreathing by a patient and for any detectors or sensors, whether basedon mechanical, electronic or other apparatus, for detecting patient flowrate as a function of time.

[0115] It will be noted on reference to FIG. 7 that relatively sharpbreaks or slope changes occur in the flow rate trace when patientinhalation effort begins and ends. Thus, although the slope of the flowrate trace may vary widely from patient to patient depending on thepatient resistance and elastance, for any given patient it will tend tobe relatively constant during inhalation and exhalation. The relativelysudden increase in flow trace slope when an inspiratory effort begins,and a corresponding sudden decrease in slope when inspiratory effortends, can be used to identify IPAP and EPAP trigger points. Since theabrupt changes in flow trace slope at the beginning and end ofinhalation correspond with the beginning and end of inspiratory effort,the flow trace slope or shape may be used to trigger the system betweenthe IPAP and EPAP state in reliance on changes in patient effort ratherthan on relatively fixed flow thresholds.

[0116] The basic algorithm for operation of this further improvedtriggering system is as follows. The input to the system is the currentflow rate at any point on the flow trace, for example point 130 in FIG.7. The flow value is differentiated to get the corresponding currentslope 132 of the flow function at time T. Of course, the slopecorresponds to the instantaneous rate of change of flow rate. The slope132 is scaled by a time factor Δt, and is then added to the current flowvalue. This gives a prediction of the flow 134 at a future time t+Δt,based on the current slope 132 of the flow trace. As noted, themagnitude of time interval Δt is determined by the selected scale factorby which slope 132 is scaled. The resulting flow prediction 134 assumes,effectively, a uniform rate of change for the flow throughout the periodΔt. It will be understood that in FIG. 7 the magnitude of time intervalsΔt is exaggerated, and is different for the end of inspiration andbeginning of inspiration changes, to facilitate clear illustration. Inpractice, Δt may be approximately 300 milliseconds, for example.

[0117] The accuracy of the predicted future flow rate value 134 willdepend upon the actual, varying slope of the flow trace during timeinterval Δt as compared to the presumed uniform slope 132. To the extentthat the flow trace during interval At deviates from slope 132, theactual flow value 136 at the end of time interval At will vary from thepredicted flow value 134. Accordingly, it may be seen that theeffectiveness of this algorithm depends upon the actual deviation in theflow trace from slope 132 during any given Δt time interval being smallexcept when the patient makes a significant change in respiratoryeffort.

[0118] To further modify the predicted flow 134, an offset factor may beadded to it to produce an offset predicted value 138. The offset isnegative during inspiration so that the predicted flow at time t+Δt willbe held below the actual flow at t+Δt except when the flow trace changesin the decreasing flow direction, such as would occur at the end ofinspiration when patient effort ceases or even reverses.

[0119] The above description refers to an end-of-inspiration trigger toEPAP. In an entirely similar fashion, as indicated generally at 140 inFIG. 7, a trigger to IPAP can be governed by prediction of a future flowrate based on differentiation of an instantaneous flow rate 139 plus anoffset factor to provide a predicted flow rate value 142. It is notedthat in this instance the offset factor will be a positive offset sinceit is an abruptly increasing flow rate that one would wish to detect. Inthis instance, when actual flow rate increases sufficiently during theAt interval to exceed predicted value 142, for example as indicated at144, the system will trigger to IPAP. Thus, at any point in the flowcycle when actual flow after a predetermined time interval At differssufficiently from the predicted, offset flow value based on flow traceslope at time T, the system is triggered to IPAP or EPAP depending uponthe direction, either positive or negative, in which actual flow variesfrom the predicted flow.

[0120] The flow rate prediction based on current flow trace slope, andthe comparison of the predicted flow rate with actual flow rate asdescribed, is repeated at a high rate to generate a continuous or nearlycontinuous stream of actual flow-to-predicted flow comparisons Forexample, in an analog system the process is continuous, while in digitalsystem the process is repeated every 10 milliseconds or faster. In theresulting locus of predicted flow value points, most actual-to-predictedflow comparisons do not result in triggering because actual flow ratedoes not deviate sufficiently from the predicted rate. Only at theinspiratory/expiratory changes does this occur. The result is atriggering method which, because it is based on flow predictions, doesnot require changes in patient inspiratory effort to achieve triggeringin synchronism with spontaneous patient respiration. Of course, thedescribed differentiation technique of flow prediction is but oneexample of a suitable flow wave shape triggering algorithm.

[0121] A schematic diagram of one analog circuit embodying elements ofthe above-described improved trigger system is shown in FIG. 8. Inputsignal 146 is estimated patient flow, although it may alternatively betotal flow. Use of estimated flow, in accordance with other aspects ofthe invention as described hereinabove, allows for further improvedtriggering if the system includes an unknown leak component. The inputsignal 146 is differentiated by inverting differentiator 148. Thediodes, switches and operational amplifier group indicated generally at150 are included for practical considerations. They form a switchablepolarity ideal diode that is used to clip the large positive derivativeof the flow trace at the beginning of inspiration and the large negativederivative at the beginning of expiration. These derivatives are usuallylarge, especially when ventilator pressure changes occur. If the largederivatives are not clipped, they can interfere with trigger circuitoperation in the early part of each respiratory phase.

[0122] The differentiated input signal is scaled by feedback resistor152 and input capacitor 154. These have a time constant equal to 300milliseconds which is a preferred delay time for the delay portion ofthe circuit. The series input resistor 156 limits high frequency noisewhich is outside the range of breathing frequencies of interest. Thedifferentiated input signal is subtracted from the flow signal in adifference amplifier 172 since the derivative circuit output is invertedand a sum is needed.

[0123] The requisite delay is produced using two fifth order, switchedcapacitor Bessel low pass filters 158. The Bessel filter has a linearphase shift with frequency which has the effect of providing a timedelay. Since the time delay depends on the order of the filter as wellas the frequency response, two filter sections are needed to provide ahigh enough cutoff frequency with the 300 millisecond time delay. The300 millisecond delay was determined experimentally. This value seemsappropriate in that the time constants for respiratory muscle activityare on the order of 50 to 100 milliseconds. Thus, a delay longer than 50to 100 milliseconds would be needed to intercept flow changes caused bychanges in respiratory muscle activity.

[0124] A 555 type timer chip 160 is set up as a 1 kHz oscillator. Alongwith the RC combination 174 at the Filter input, element 160 controlsthe cutoff frequency of the switched capacitor filters 158. Finally, acomparator 162 with hysteresis, provided by input resistor 176 andfeedback resistors 166 and 164, changes its state based on thedifference between the input flow rate signal and the processed flowratesignal. There is a fixed hysteresis during the inspiratory phase causedby resistor 166. The hysteresis characteristic of the described circuitprovides the negative offset during inspiration and the positive offsetduring expiration mentioned above.

[0125] During the expiratory phase, the hysteresis is initially greaterthan that during inspiration and is based on the magnitude of the flowsignal as set by the current through resistor 164 and diode 178. Theinverting amplifier 180 has a gain of 10 so that it saturates when theflow signal is greater than 0.5 volts, which corresponds to a flow of 30liters per minute. The hysteresis is therefore almost doubled during theinitial part of the exhalation phase by virtue of the look resistor 164in parallel with the 82 k resistor 166, in contrast with the 82 kresistor 166 acting alone during inspiration. That is, during exhalationthe output of invertor 180 is positive. Therefore, diode 178 is forwardbiased and supplies current through resistor 164 in addition to thatsupplied by resistor 176.

[0126] Once the flow signal drops below the 0.5 volt level, thehysteresis decreases linearly with flow rate to be the same as thatduring inspiration. This feature prevents premature triggering toinspiration while gradually increasing the sensitivity of the systemtoward the end of exhalation. The hysteresis is somewhat dependent onthe flow signal since it is supplied through a resistor 164. With thecomponents described (82 k feedback resistor 166, 1.5 k input resistor156) and a +5 volt system power supply (not shown), the hysteresis isabout 90 millivolts at zero flow. This corresponds with a trigger offsetof 90 cc per second above the instantaneous flow signal.

[0127]FIG. 9 illustrates operation of the flow trace shape triggercircuit with the heavy line 168 indicating the flow signal at the inputand the lighter line 170 indicating the processed flow signal at the +input to the system comparator 162. Line 170 shows hysteresis decreasingas flow approaches zero during exhalation.

[0128] The triggering algorithm described immediately above is referredto as shape triggering (i.e. relying on changes in the shape or slope ofthe flow trace) to distinguish it from the flow triggering algorithmdescribed earlier. FIGS. 10A and 10B illustrate the behavior of twoventilator triggering algorithms when used in a simulation of aspontaneously breathing patient with a large respiratory time constant.In each of FIGS. 10A and 10B, the uppermost trace represents estimatedpatient flow, the center trace represents patient respiratory effort,and the bottom trace represents ventilator pressure generated at themask which interfaces with the patient airway. The letter I indicatespatient inspiratory effort in each respiratory cycle.

[0129] As may be seen, FIG. 10A illustrates the sorts of triggeringproblems which were described hereinabove and which have been known tooccur with actual patients. Patient inspiratory effort does notconsistently trigger the ventilator, as shown by the asynchrony betweenpatient effort, the estimated patient flow, and pressure generated atthe mask. FIG. 10B, by contrast, illustrates ventilator response to thesame simulated patient using the above-described shape triggeringalgorithm, but with all other settings unchanged. A comparison of FIGS.10A with 10B reveals the improved synchronization of delivered pressurewith spontaneous patient effort.

[0130] It is noted that the simulated patient on which FIGS. 10A and 10Bare based exhibits significant exhalatory flow at the beginning of aninspiratory effort. From FIG. 10B, it may be clearly seen that the shapetriggering algorithm triggers appropriately at the beginning and end ofeach and every inspiratory effort even though the inspiratory effortbegins while there is still exhalatory flow. Thus, the shape triggeringalgorithm achieves the objective of synchronizing ventilator triggeringwith changes in patient effort.

[0131]FIGS. 11A and 11B are similar to FIGS. 10A and 10B, respectively,but the simulated patient has a shorter respiratory system timeconstant. FIG. 11A illustrates, from top to bottom, estimated patientflow, patient respiratory effort, and pressure at the mask for a givertriggering algorithm, and FIG. 11B illustrates the same parameters forthe above-described flow trace shape triggering algorithm. In FIG. 11Ait is clear that even with a substantial drop in flow when patienteffort stops, the system does not trigger to EPAP until flow drops belowabout 25% of peak flow. Thus, the end of the inspiratory phase is notsynchronized with the end of patient inspiratory effort. By contrast,the shape triggering algorithm of FIG. 11B produces, for the samepatient, consistently synchronized triggering at the end of patientinspiratory effort.

[0132] As may be appreciated from the above description, a ventilatorcan be triggered on the basis of changes in the shape of the flow signalwhich correspond to changes in patient respiratory effort. Thistriggering algorithm solves several vexing problems which have limitedthe utility of some prior triggering algorithms such as standardpressure support triggering. It is also to be noted that the describedflow signal shape triggering algorithm does not rely on any calculationof system leakage, although the shape triggering algorithm can operatemore reliably with an estimate of system leakage than without it.

[0133] Although the described shape triggering algorithm works totrigger both to IPAP and to EPAP, it tends to work best for triggeringto EPAP because normally a much larger and more abrupt change in patientrespiratory effort occurs at the end of inspiration than at thebeginning.

[0134] It will be noted further that the described shape triggeringalgorithm may be used in parallel with the earlier described flowtriggering algorithm to provide a dual triggering system in which asecondary triggering algorithm will function in the event the primaryalgorithm malfunctions. This sort of dual triggering system can operateon a continuing breath-to-breath basis such that at any desired triggerpoint the secondary triggering algorithm will trigger the ventilator ifthe primary triggering algorithm does not.

[0135] Additional improvements in leak compensation techniques are alsocontemplated by this invention. As noted in the above descriptionconcerning leak compensation, the algorithm described there relies ontwo requirements as follows: (1) the patient's inhaled and exhaledvolumes over time are the same, (and indeed if the patient's rest volumejust prior to the beginning of inspiration is the same from breath tobreath, the inhaled and exhaled volumes for each individual breath willalso be the same); (2) when the patient is inhaling, total flow isgreater than leak flow and when the patient is exhaling total flow isless than leak flow.

[0136] In the above description relating to leakage, only total patientcircuit flow is actually measured even though this flow is made up oftwo components, patient flow and leakage flow. Further, in the abovedescription concerning leakage the leak is not estimated as a functionof pressure. Rather, an average leak is calculated by integrating totalflow. Since patient inhalation volume and exhalation volume areessentially the same, the average flow over a complete respiratory cycleis generally equal to the average leak per cycle.

[0137] It is to be noted further that the leak component itself can havetwo components, namely the known or intended leak of an exhalation portand an unknown leak component resulting from one or more inadvertentleaks such as leakage across a mask seal or at a tubing connection. Theunknown or inadvertent leak component would be quite difficult todetermine exactly as it can be a function of both pressure and time.Therefore, to the extent it is necessary to determine this leakcomponent, its value is estimated; however, significant leak managementimprovements can also be developed without separating the leak into itsintended and inadvertent components.

[0138] Regardless of whether or not the overall leak can be characterizeas a function of time or pressure, the objective is still to adjust theestimate of leakage so that its average value is the same as the averagevalue of the true leak over integral respiratory cycles. Of course, theaverage of the true leak can be obtained as the average of total flowsince the patient flow component of average total flow is zero.

[0139] If one assumes the leak is a function of pressure, such as theleak from a WHISPER SWIVEL (tm) connector, we then can assume the leakwill likely be less at EPAP than it is at IPAP, with an average valuecorresponding to a pressure between EPAP and IPAP. That is, for a leakwhich is a function of pressure, the leakage rate at a lower pressure isless than the leakage rate at a higher pressure. This characteristic ofleak flow can be utilized in algorithms other than that describedhereinabove to compensate for system leakage.

[0140] In all systems described herein, the purpose of determiningleakage is to thereby determine patient flow in an open system. Ofcourse, a closed system ideally has no leaks, but in practice may haveinadvertent leaks; however, in an open system gas flow from the system,whether by a regulated leak at an exhaust point or through inadvertentleaks from tubing connections and seal interfaces, will constitute asignificant part of the total system flow. Accordingly, leakage in thepatient circuit of an open system can be calculated by using the totalflow as detected by a pneumotach. This parameter, which we refer to asraw flow, includes all flow leaving or entering the gas flow source(eg., a blower) via the patient circuit. Therefore, raw flow includesboth patient flow and system leakage. The purpose of the described leakmanagement algorithm is to separate raw flow into patient flow andleakage components.

[0141] It is not necessary to directly measure either the patient flowor leakage component of raw flow, but consequently the results of theleakage calculations are to be regarded as estimated values. One mode ofleak management is discussed hereinabove with reference to FIGS. 1 to 5.The following algorithm description relates to alternative approaches toleak management.

[0142] For a non-pressure dependent leak, one first determines whenpatient inhalation begins. We refer to this point in the patient'srespiratory cycle as a the breath trigger and use it as a referencepoint to compare patient inspiratory and expiratory volumes. By usingthe beginning of inspiration as a breath trigger point, the leakagecalculation is done with the patient's beginning lung volume atapproximately the same level for every calculation.

[0143] The breath trigger reference point is identified by satisfactionof two conditions as follows: (1) raw flow greater than average leakrate; and (2) patient is ready to inhale. Concerning the first of theseconditions, the comparison of raw flow with average leak rate revealswhether the patient is inhaling or exhaling because, as noted above, thenature of the raw flow component is such that raw flow exceeds systemleakage during patient inspiration and is less than system leakageduring patient expiration.

[0144] The second condition, that of determining whether the patient isready to inhale, is assessed by reference to either of two additionalconditions as follows: (1) the patient has exhaled a predeterminedfraction of the inhaled volume, for example ¼ of inhalation volume; or(2) the patient has exhaled for more than a predetermined time, forexample 300 milliseconds. An alternative statement of the first of theseconditions is that the patient's exhaled volume is at least asignificant fraction of inspiratory volume. As to the second condition,an exhalation time of more than 300 milliseconds can be detected bycomparison of raw flow with estimated average leakage rate since, asnoted above, raw flow is less than estimated average leak rate duringexhalation. If raw flow thus is less than the average leak rate forlonger than 300 milliseconds, the condition is met. If either of theconditions 1 or 2 immediately above is met the patient is ready toinhale. Therefore, if either of the conditions is met and in additionthe raw flow is greater than average leak rate, the conditions for abreath trigger are satisfied and the described conditions thus can beused to trigger the ventilator for an inspiratory cycle.

[0145] The seeming contradiction of requiring raw flow greater thanaverage leak rate to satisfy one breath trigger condition, and raw flowless than average leak rate as one parameter that can satisfy the secondbreath trigger condition is explained as follows. The condition that rawflow is greater than average leak rate merely indicates that the patienthas begun an inhalation cycle; that is, the patient has exerted theinitial inspiratory effort that is revealed as raw flow exceedingaverage system leakage. However, as discussed hereinabove, since the rawflow signal can and does cross the average leak rate signal due tospurious noise in the raw flow signal, not all incidents of raw flowbeing greater than average leakage will denote the beginning of patientinspiratory effort. In fact, noise in the raw flow signal may typicallyresult in multiple crossings between the raw flow signal and the averageleak signal. Especially troublesome in this regard is noise causing amomentary − to + transition while the actual major condition is the flowcrossing zero due to a true inspiratory to expiratory transition.

[0146] In order to avoid multiple triggers at these points of signalcrossings that do not denote initial patient inspiratory effort, theadditional ready-to-inhale condition is imposed. Under this condition,the raw flow greater than average leak condition is validated only ifone of the two conditions indicating the patient's respiration was justpreviously in an exhalatory state is satisfied. In the case of thesecond of these-conditions, the 300 millisecond duration of raw flowless than average leakage corresponds to patient exhalation. Thus, thetime element plays a role in the described breath trigger algorithm andmust be understood. Raw flow less than average leakage indicatesexhalation which then provides a ready-to-inhale output. A subsequentreversal of raw flow to a value greater than average leakage indicatesthe detection of patient inspiratory effort. The breath trigger thusconsistently initiates an inspiratory cycle in synchronism withspontaneous patient respiration.

[0147] When the breath trigger occurs, the ready-to-inhale parameter isreset to await satisfaction of one of the two conditions specified abovethat provides a ready-to-inhale validation. The breath triggeringprocess is then repeated upon occurrence of the next patient inspiratoryeffort when raw flow again exceeds average system leakage.

[0148] The time interval between breath triggers from one breath to thenext can be captured and the raw flow signal integrated over that timeinterval to find the raw volume or total volume for each breath. The rawvolume may then be divided by the time between breath triggers, forexample the time interval between the next successive pair of breathtriggers, to determine a recent time rate of leakage.

[0149] It is noted that the patient's respiration from one breathtrigger to the next, as represented by the integration of the raw flowsignal, rises from zero volume to a maximum value comprised ofinspiratory volume and leakage volume during inspiration. As therespiratory cycle continues through expiration, the continuedintegration of negative flow reduces the raw volume parameterprogressively as the exhaled volume is subtracted. Thus, at the nextbreath trigger, the raw volume parameter from the prior breath is equalto only the leakage volume and any change in patient resting volume;however, this latter value generally is assumed to be zero.

[0150] Thus, as the raw volume parameter is really only leak volume,dividing that value by the time duration of the breath provides anestimate of the leak rate for the most recent breath. By averaging therecent leak rate over time an average leak rate can be determined toprovide a more stable signal. Among other benefits, this reduces anyeffects of breath-to-breath volume variation resulting frombreath-to-breath changes in patient resting volume. The number ofrespiratory cycles over which the recent leak may be averaged todetermine average leak rate may be anywhere between one and infinity,depending upon the desired balance between signal stability and rapiderror correction. Longer term averaging results in greater stability butslower error correction. Shorter term averaging provides less stabilitybut quicker error correction.

[0151] Finally, it is to be noted that the average leak rate whosecalculation is described here was used initially in this algorithm inone of the conditions for initiating a breath trigger. The parametersused in the conditions for initiating a breath trigger thus can becontinually updated, used to satisfy the breath trigger conditions, andthen the successive breath triggers are used in turn to again updatethese parameters.

[0152] A related approach to leakage analysis involves the use of a leakcomponent which is, or is assumed to be, a function of patient circuitpressure. This sort of leak analysis can be useful in such ventilationregimens as, for example, proportional assist ventilation such asdescribed in U.S. Pat. No. 5,107,830. It is important in proportionalassist ventilation to know both the average leak rate and theinstantaneous leak for the system. However, proportional assistventilation involves variation of pressure according to the level ofventilation assistance required by the patient from moment to moment.Since at least some components of system leakage can be a function ofpressure, those components will also vary more or less in synchronismwith varying patient ventilation needs.

[0153] Thus, an alternative algorithm for leak analysis that can be usedin a proportional assist ventilation system and in other systems asdeemed suitable would be based on the requirement that patient flowequals raw flow, as defined hereinabove, less any known leak componentand any pressure dependent leak component. Known leakage may be anyleak, intentional or otherwise, having known flow characteristics, forexample the leakage through a WHISPER SWIVEL (tm) or through a plateauexhalation valve. That is, the known leak is not necessarily a fixedleak, but can also be a function of pressure. Although such leaks ofknown characteristics could be lumped together with leaks having unknowncharacteristics but assumed to be functions of pressure, the followinganalysis can be carried out with leaks of known characteristics includedor excluded from the leakage calculation described. The pressuredependent leak, whether constituted of total leakage or only a componentof total leakage, is calculated as a function of system pressure fromanalysis of the raw flow signal. The function by which the pressuredependent leak component is related to pressure is that for any orificedefining a pressure drop between a pressurized system and a lowerpressure surrounding environment, the flow through the orifice will beproportional to the square root of the pressure difference across theorifice multiplied by a constant K which characterizes the mechanicalfeatures of the orifice itself, that is its size, surface smoothness,and so forth.

[0154] To find the pressure dependent leak component, it is necessaryonly to characterize one hypothetical orifice as representing the sourceof all pressure dependent leakage. To find the hypothetical orifice, thebreath trigger defined above is utilized to mark the beginning and endof patient respiratory cycles, and the average patient circuit pressureis measured for each breath. The average leak rate, as calculatedhereinabove, is multiplied by the time duration of the previous breathto determine pressure dependent leak volume on a per-breath basis. Ifthe pressure dependent leak component is separated from the known leakcomponent, then the known leak component is also separated from averageleak rate before the average leak rate is used to determine pressuredependent leak volume. The pressure dependent leak volume is thendivided by the square root of the average pressure for the prior breathto give the corresponding orifice characteristic K. Over the nextbreath, the pressure dependent leak is found by multiplying thecalculated orifice characteristic K by the square root of instantaneouspressure. This gives the pressure dependent leak rate as a function ofpressure throughout the breath. This is an important parameter in suchventilation regimens as the above characterized proportional assistventilation. More generally, however, it is important for the therapistto have reliable information on system leakage. The pressure dependentleak rate therefore can be useful for characterizing leakage in avariety ventilation systems.

[0155] For either of the last described leak calculations, recoveryroutines are desired to deal with those instances when a breath triggerdoes not occur when it should due to failure of the raw flow signal tocross the average leak rate signal as the patient breaths. When thesystem leak changes, the raw flow signal will increase if the leakincreases or decrease if the leak decreases. It is possible for thisincrease or decrease to be large enough that the raw flow signal nevercrosses the average leak rate, and when this occurs there will be nobreath trigger and thus no new or updated leak calculations. To recoverfrom such an incident, an algorithm is needed to begin the leakagecalculations again and thereby bring the calculated average leak rateback into line with actual system functioning. Initiation of therecovery algorithm is determined to be necessary if any one of fourknown physiological events occurs as follows: (1) exhalation continuesfor more than 5 seconds; (2) inhalation continues for more than 5seconds; (3) inspiratory tidal volume is greater than 5 liters; or (4)expiratory tidal volume is less than −1 liter.

[0156] The ventilator system should include elements for monitoringoperation to detect occurrence of any one of these events. When one suchevent is detected, rediscovery of the leak is initiated by changing theaverage leak rate to the raw flow signal by gradually adding apercentage of the difference between raw flow and average leak to theaverage leak rate. Preferably, the percentage of difference added shouldbe selected to provide a time constant of approximately one second.

[0157] As the gradual increase in average leak rate in accordance withthe recovery algorithm causes the average leak rate value to approachraw flow value, the increasing average leak rate ultimately will cause abreath trigger in accordance with one of the triggering algorithms asdescribed above, and the repetitive leak calculations will then resumewith each respiratory cycle.

[0158] From the above leak management algorithm it may be seen that thealgorithm may also be based on comparison of estimated patient flow to azero value, and calculate only an unknown leak as a function ofpressure. In a manner similar to the leak management algorithm abovecharacterized, a breath indicator or trigger is employed as a marker forthe start of patient inspiration. The breath trigger occurs when twoconditions are met, namely: (1) estimated patient flow is greater than0; and (2) patient is ready to inhale. The ready-to-inhale condition issatisfied when: (1) the patient has exhaled a specified portion (eg.25%) of inspiratory volume; (2) estimated patient flow is less than 0;and (3) inspiratory volume is greater than a constant represented by 0or a small positive value.

[0159] As with the above-described algorithms, the conditions requiringinspiratory volume to be both less than a percentage of expiratoryvolume and greater than a small positive value are not contradictory asthese measurements are taken at different times. First, theready-to-inhale condition must be satisfied by the specified parameters,including detection of an estimated patient flow less than zero sincethe last inhalation. Then, a breath trigger can occur when estimatedpatient flow goes positive since estimated patient flow greater than 0indicates initial patient inhalatory effort.

[0160] From one breath trigger to the next, raw flow is integrated andthe square root of pressure is also integrated. Then at the next breathtrigger the unknown orifice is calculated by dividing the value obtainedfrom raw flow integration by the value obtained from integration of thesquare root of pressure to provide the characteristic K for the unknownorifice. As noted above, the unknow orifice is a hypothetical leaksource intended to describe the characteristic of leakage from allsources of pressure dependent leak Once the unknown orificecharacteristic has been calculated, it is multiplied by the square rootof patient circuit pressure at any time during the following breath toprovide a measure of the instantaneous unknown leak. Estimated patientflow is then found by subtracting the instantaneous unknown leak fromthe raw flow.

[0161] In both of the algorithms which treat the system leak as the gasflowing through a hypothetical orifice of characteristic K, the constantK is calculated in essentially the same way. Although the equation forthe calculation is derived differently for the two approaches, theresult is the same, as follows. $\begin{matrix}{K = \frac{\int_{0}^{T}{\left( {{total}\quad {flow}} \right){t}}}{\int_{0}^{T}{\left( \sqrt{Pressure} \right){t}}}} & {{eq}.\quad 1}\end{matrix}$

[0162] In cases where the system also includes a known leak componentsuch as a leak resulting from use of a plateau valve or a WHISPER SWIVEL(tm) as described above, the calculations can be modified to providegreater accuracy by subtracting the known leak component from raw flowprior to integration of raw flow to determine the unknown orifice, andalso prior to utilizing raw flow to determine estimated patient flow. Inthis case, the integral of raw flow minus the known leak, divided by theintegral of the square root of patient circuit pressure will provide theparameter K characterizing the unknown orifice attributable to only theunknown leak. That orifice characteristic K can then be multipliedby-the square root of patient circuit pressure during the subsequentbreath to provide a measure of the unknown leak component. Estimatedpatient flow at any time then would be found by subtracting both theknown leak and the calculated unknown leak from the raw flow signal.

[0163] The modified algorithm preferably also includes a recoveryroutine for essentially the same reasons as earlier stated. A breathtrigger for this modified algorithm depends upon estimated patient flowcrossing the zero value as the patient breaths. When the leak valuechanges, both raw flow and estimated patient flow will increase if theleak is increasing, or decrease if the leak is decreasing. It ispossible for such an increase or decrease to be large enough thatestimated patient flow will never cross the 0 value. When this occurs nobreath trigger occurs and there will be no new or updated leakcalculations unless a recovery routine is performed.

[0164] To correct this condition, a recovery algorithm may be utilizedto detect when leakage calculation is out of control according to twolimiting parameters as follows: (1) inspiratory tidal volume is greaterthan a physiological value such as 4.5 liters; or (2) inspiratory tidalvolume is less than a non-physiological value such as −1 liter. Wheneither of these non-physiological events occurs, the leakage calculationis out of control and the system leak must be rediscovered byartificially inducing a breath trigger. Thus, the system would requireflow detecting apparatus to sense the specified limiting values ofinspiratory tidal volume and provide a breath trigger reference point.This restarts system operation essentially in the same manner asdescribed above with reference to other algorithms, and thus initiates anew leak calculation. The data used for these continued calculations caninclude data from after the change in leak magnitude so that theestimate of the unknown leak is improved. Even if the new unknown leakvalue is incorrect, continuing repeated calculations also will be basedon data from after the change in leak magnitude. Thus, the calculatedestimate of unknown leak will continue to quickly approach reliablevalues.

[0165]FIG. 12 illustrates the described leak management algorithm inaction, with intentional introduction of a very large leak. In FIG. 12,the four traces represent, from top to bottom, pressure, estimatepatient flow, estimated inspiratory volume, and estimated leak (combinedknown and unknown leakage). The X axis is progressing time moving fromleft to right.

[0166] As can be seen from FIG. 12, for the first three respiratorycycles the algorithm is able to determine leakage as a function ofpressure. At the start of each inspiration an unknown orificecharacteristic K is calculated. Points 1 and 2 on the estimated leaktrace show the transitions between one unknown orifice characteristic Kand the next, which is an updated unknown orifice characteristic. Point2 on the estimated leak trace shows a much smoother transition betweenold and new unknown orifice characteristics than does point 1,thus-indicating that point 2 is a transition between unknown orificecharacteristics of essentially the same value.

[0167] To test algorithm response to leakage, a very large leak wasintroduced as indicated at point 0 on the estimated patient flow trace.Predictably, the estimated patient flow and estimated inspiratory volumerise off scale very quickly. At time T, estimated inspiratory volume hasreached 4.5 liters and a breath trigger thus is forced. The resultingcalculation estimates the leak to be a very high value which, in theestimated leak portion of FIG. 12 is off scale. The resulting estimatedpatient flow is approximately 50 liters per minute where in fact itshould be zero. Tidal volume continues to rise rapidly as indicated atpoint 3 and would ultimately force another breath trigger when itreached the 4.5 liter limit. The large system leak is removed at timeT′, however, so the rise in tidal volume ceases.

[0168] With the leak removed from the system, estimated patient flow andestimated inspiratory volume rapidly drop off scale. Thus, wheninspiratory tidal volume reaches the recovery limit of −1 liter, anotherbreath trigger is forced thus leading to a new leak calculation and aresulting estimated leak lower than the previous leak calculation. Thisbrings estimated patient flow back onto scale. Although the new leakcalculation is better than the previous one, it is not by any meansperfect. With estimated patient flow based on this new unknown leakorifice characteristic, tidal volume may once again drop below the −1liter limit such as indicated at point 4 thus forcing another breathtrigger. Once again a new unknown orifice calculation is performed andused for an unknown leak calculation which results in a very good value,thus returning the system to normal ranges of operation. The system thuswas reliable in recovery from even very large changes in leak magnitude.The much smaller changes in leak magnitude that are more commonlyencountered are readily handled under this algorithm by either thenormal calculation or the recovery calculation.

[0169] Of course, we have contemplated various alternative and modifiedembodiments of the invention of which the above described areexemplary-as the presently contemplated best modes for carrying out theinvention. Such alternative embodiments would also surely occur toothers skilled in the art, once apprised of our invention. Accordingly,it is intended that the invention be construed broadly and limited onlyby the scope of the claims appended hereto.

We claim:
 1. A method of detecting changes in the respiratory state of apatient between inspiration and expiration comprising the steps of:monitoring the respiratory gas flow of such a patient; from saidmonitoring, determining a patient respiratory gas flow characteristic atselected time intervals; comparing selected pairs of said respiratorygas flow characteristics; and from said comparing, identifying changesin the patient's respiratory state between inspiration and expiration.2. The method as set forth in claim 1 wherein said comparing stepincludes comparing a later determined respiratory gas flowcharacteristic with a fraction of an earlier determined respiratory gasflow characteristic.
 3. The method as set forth in claim 2 wherein saidrespiratory gas flow characteristic is the patient's respiratory gasflow rate.
 4. The method as set forth in claim 1 wherein each saidselected pair of respiratory gas flow characteristics includes apredicted value of the patient's respiratory gas flow rate at a selectedtime, and the patient's actual respiratory gas flow rate at saidselected time.
 5. The method as set forth in claim 4 including theadditional step of taking the derivative of a known patient respiratorygas flow rate and processing said derivative to determine said predictedvalue of the patient's respiratory gas flow rate.
 6. The method as setforth in claim 5 wherein said processing step includes scaling saidderivative with a time factor and modifyin the scaled derivative with anoffset factor to determine said predicted value of the patient'srespiratory gas flow rate.
 7. A method of determining the leak componentof gas flow in a respiratory gas supply system for supplying respiratorygas to a patient comprising the steps of: determining the beginning andend of an integral number of patient breaths; integrating patientrespiratory gas flow rate throughout at least some of said patientbreaths to determine total respiratory gas volume for said at least someof said patient breaths; and dividing said total respiratory gas volumefor said at least some of said patient breaths by the time duration ofthe respective patien breaths to determine a leakage rate for said atleast some of said patient breaths.
 8. The method as set forth in claim7 including the additional step of averaging said leakage rates of saidat least some of said patient breaths to determine an average rate ofrespiratory gas leakage from such a respiratory gas supply system.
 9. Amethod of determining the rate of leakage from a respiratory gas supplysystem which supplies respiratory gas flow to a patient at specifiedflow rates and pressures comprising the steps of determining said rateof leakage as a leak which is proportional to the square root ofpressure by a constant of proportionality K, where $\begin{matrix}{{K = \frac{\int_{0}^{T}{\left( {{total}\quad {flow}} \right){\quad t}}}{\int_{0}^{T}{\left( \sqrt{Pressure} \right){\quad t}}}},} & {{eq}.\quad 1}\end{matrix}$

the numerator of eq. 1 representing average leak flow and thedenominator of eq. 1 representing the relationship between the leak rateand pressure.